New areas of medical study and new clinical applications involving the use of 500 KHz-300 MHz frequency ultrasound imaging are constantly being developed. Ultrasound images made at the high end of this frequency range have spatial resolutions that approach the wavelength of the ultrasound energy, e.g., 20 microns for a 75 MHz ultrasound signal in water. Initial clinical applications of high frequency ultrasound include imaging the eye, the vasculature, the skin, and cartilage. Such imaging may be used, for example, to determine the vertical growth phase of skin cancers, to distinguish between cancerous tissue and fat in the breast, and to determine quantitative information about the structure of atherosclerotic plaque in arteries.
Future improvements in ultrasound image quality will require the fabrication of ultrasonic transducer arrays using designs and fabrication techniques not heretofore available. More particularly, transducer arrays manufactured with current transducer fabrication technology have limited spatial resolution, restricted scan slice thickness, inadequate phase correction capability, and primitive beam steering for volumetric scanning. To overcome these limitations, the next generation of ultrasonic transducer arrays will need to be multi-dimensional and operate over a broad range of frequencies.
Ultrasound imaging arrays having a 2-D (N.times.M) configuration are the subject of much research and development due to their potential for overcoming some of the above-described limitations of known one-dimensional (N.times.1) linear arrays. Unfortunately, rapid development and commercialization of 2-D ultrasound imaging arrays has been hampered by difficulties in fabricating transducer elements with small dimensions and low electrical impedance.
Once 2-D ultrasonic imaging arrays having improved resonant frequency, sensitivity and other operating characteristics are developed, it is anticipated a number of ultrasound applications will become available. First, focusing could be performed in an elevation plane that is perpendicular to the primary imaging plane at slice thicknesses and image resolutions not currently available. Second, cross axis phase aberration caused by differences in ultrasonic propagation velocity through different tissue types could be corrected through the use of 2-D imaging arrays. Third, 2-D arrays with improved sensitivity and resolution will allow true volumetric imaging of structures that are too small to be imaged with current technology.
Another area of current interest, high-intensity focused ultrasound (HIFU), has significant potential for use in therapeutic ultrasound applications including noninvasive myocardial ablation, drug delivery, drug activation, ultrasound surgery and hyperthermia cancer therapy. Ideally, HIFU therapies would be performed while simultaneously viewing the area being treated. For example, for therapy, high power sound bursts are delivered at one frequency, while for imaging, a different frequency may be desirable to provide images with sufficient resolution.
Unfortunately, known ultrasonic imaging systems do not typically permit such dual application of ultrasound with a single transducer array. Instead, with current systems, the body region to be treated is generally imaged with a first transducer, and then the HIFU therapy is administered with a second transducer. Introduction of an ultrasound transducer into certain body regions can be a relatively lengthy, e.g., 45 minutes, and risky procedure. Also, appropriate placement of the transducer delivering the HIFU therapy is a challenge given the absence of contemporaneous imaging information.
In an attempt to address this limitation with known ultrasonic imaging systems, experiments have been conducted using broadband ultrasonic transducers, i.e., ceramic and composite-based transducers having an upper frequency that is about 1.6 times the center frequency and a lower frequency that is about 0.4 times the center frequency. By controlling the frequency content of the drive signal, the transducer can be operated to transmit and receive ultrasound near the opposite ends of the transducer's frequency range. However, broadband ultrasonic transducers operated in this manner have a serious shortcoming due to different characteristics of therapy and imaging ultrasound transducers. A sharp resonance is required for improved efficiency for therapy, while a broad bandwidth is required for effective imaging. In addition, in some circumstances it is desirable to provide ultrasonic energy at two frequencies that are spaced farther apart than is achievable with known broadband transducers.
Sheljaskov et al., in the article A Phased Array Antenna for Simultaneous HIFU Therapy and Sonography, Proceedings of the 1996 Ultrasound Synopsium, pages 1527-1530, describe an ultrasonic transducer capable of generating ultrasonic energy with the same transducer at 1.7 MHz and 5.5 MHz. The transducer features two piezoceramic layers stacked one on top of the other, and one matching layer. One of the piezoelectric layers is divided into three separately wired sections. The piezoceramic layer divided into three separate sections may be operated independently of the other layer to produce the 5.5 MHz signal. The 1.7 MHz signal is created by operating the entire transducer as a single unit. Thus, the separate piezoelectric sections of the transducer necessarily acoustically communicate with each other. The transducer apparently cannot be operated to provide the 1.7 MHz signal at exactly the same time it is providing the 5.5 MHz signal because the same piezoelectric ceramic is required to produce both the high and low frequency ultrasonic energy. Thus, when Sheljaskov et al. indicate their transducer provides both signals "simultaneously," it is believed they use this term loosely. In addition, it is believed this transducer faces the same limitations as other prior art broadband transducers described above, i.e., non-optimum design for two mutually exclusive uses.
Thus, a need clearly exists for an ultrasonic imaging system for providing multiple frequencies of ultrasonic energy at frequencies higher than those achievable with known imaging systems. In addition, certain ultrasound applications require higher signal-to-noise ratios, and hence resolutions, than are achievable with known dual-frequency imaging systems.
As used herein, the term "1-D array" refers to an array having (N.times.1) discrete transducer elements, the term "2-D array" refers to an array having (N.times.M) discrete transducer elements where N and M are equal or nearly equal in number, and the term "1.5-D array" refers to an array having (N.times.M) discrete transducer elements where &gt;&gt;M, e.g., where N=128 and M=3.